Radiofrequency (RF) ablation of cardiac and other tissue is a well-known method for creating thermal injury lesions at the tip of an electrode. Radiofrequency current is delivered between a skin (ground) patch and the electrode, or between two electrodes. Electrical resistance at the electrode-tissue interface results in direct resistive heating of a small area, the size of which depends upon the size of the electrode, electrode tissue contact area, and current (density). Further tissue heating results from conduction of heat within the tissue to a larger zone. Tissue heated beyond a threshold of approximately 50-55 degrees C. is irreversibly injured (ablated).
Resistive heating is caused by energy absorption due to electrical resistance. Energy absorption is related to the square of current density and inversely with tissue conductivity. Current density varies with contact area conductivity, voltage and inversely with the square of the distance from the ablating electrode. Therefore, energy absorption varies with conductivity, the square of applied voltage, and inversely with the fourth power of the distance from the electrode. Resistive heating, therefore, is most heavily influenced by distance, and penetrates a very small distance from the ablating electrode. The rest of the lesion is created by thermal conduction from the area of resistive heating. This imposes a limit on the size of ablation lesions that can be delivered from a surface electrode.
Theoretical methods to increase lesion size would include increasing electrode size, increasing the area of electrode contact with tissue, increasing tissue conductivity and penetrating the tissue to achieve greater depth and increase the area of contact, and delivering RF until maximal lesion size has been achieved (60-90 seconds for full maturation).
The electrode can be introduced to the tissue of interest directly (for superficial/skin structures), surgically, endoscopically, laparoscopically or using percutaneous transvascular (catheter-based) access. Catheter ablation is a well-described and commonly performed method by which many cardiac arrhythmias are treated.
Catheter ablation is sometimes limited by insufficient lesion size. Ablation of tissue from an endovascular approach results not only in heating of tissue, but heating of the electrode. When the electrode reaches critical temperatures, denaturation of blood proteins causes coagulum formation. Impedance can then rise and limit current delivery. Within tissue, overheating can cause steam bubble formation (steam “pops”) with risk of uncontrolled tissue destruction or undesirable perforation of bodily structures. In cardiac ablation, clinical success is sometimes hampered by inadequate lesion depth and transverse diameter even when using catheters with active cooling of the tip. Theoretical solutions have included increasing the electrode size (increasing contact surface and increasing convective cooling by blood flow), improving electrode-tissue contact, actively cooling the electrode with fluid infusion, changing the material composition of the electrode to improve current delivery to tissue, and pulsing current delivery to allow intermittent cooling.
Conventional catheters are equipped to measure temperature at their distal sections which are adapted for contact with tissue. Typically, these catheters include a thermocouple wire pair 80 and 82 that extend from the control handle, through the catheter shaft and into the distal section where a “hot” or temperature measuring junction H of the wire pair is positioned. As shown in FIG. 3B, the wire pair 80 and 82 are typically stripped of insulating cover 85, twisted, soldered and potted in a blind hole 86 formed in a distal tip electrode 84, as known in the art. However, because the wire pair 80 and 82 are stripped, twisted and normally bent back on itself with a U-turn 88, the precise location of the distal end “hot” junction H within the blind hole 86 or relative to the tip electrode 84 is not known, even though the depth of the blind hole 86 may be known. An alternate method of assembly includes laser stripping the insulation of a small area of the bifilar wire and then soldering the wires together in that location. Both methods of forming a thermocouple junction are limited in their ability to form a junction at the very bottom of a blind hole. The exact location of the thermocouple junction is the proximal location where the wires are electrically connected. Thus, the location where the temperature is actually measured is by design limited to at least the axial length of the solder joint. For certain types of catheters, it is desirable to have location accuracy of the hot junction down to at least tenths of millimeters and place the thermocouple junction as deep into the hole as possible. Thus, there is a desire for a catheter with a temperature sensor where a more precise location is known of where the temperature is being sensed. Moreover, the conventional thermocouple wire pair has an awkward profile, as shown in FIG. 3C, which can often split or tear between the two wires.
Thus, there is a desire for a catheter with a temperature sensor with a more precise location of its temperature sensing element. In particular, there is a desire for a catheter with a thermocouple wire pair having a more precise location of its “hot” junction, an improved profile, a more durable construction, and an easier method of assembly. Where space is always a constraint within a catheter, there is a further desire for a catheter with a thinner thermocouple wire pair.